U.S. patent number 6,263,221 [Application Number 08/799,206] was granted by the patent office on 2001-07-17 for quantitative analyses of biological tissue using phase modulation spectroscopy.
This patent grant is currently assigned to Non-Invasive Technology. Invention is credited to Britton Chance, Vasilis Ntziachristos.
United States Patent |
6,263,221 |
Chance , et al. |
July 17, 2001 |
Quantitative analyses of biological tissue using phase modulation
spectroscopy
Abstract
A spectroscopic system for quantifying in vivo concentration of
an absorptive pigment in biological tissue includes an oscillator
for generating a first carrier waveform of a first frequency on the
order of 10.sup.8 Hz, a light source for generating light of a
selected wavelengths modulated by the carrier waveform, and a
detector for detecting radiation that has migrated over photon
migration paths in the tissue from an input port to a detection
port spaced several centimeters apart. The wavelength is sensitive
to concentration of an absorptive pigment present in the tissue. A
phase detector compares the detected radiation with the introduced
radiation and determines therefrom the phase shift of the detected
radiation. A processor quantifies the concentration of the
absorptive pigment by calculating a value of the absorption
coefficient.
Inventors: |
Chance; Britton (Marathon,
FL), Ntziachristos; Vasilis (Philadelphia, PA) |
Assignee: |
Non-Invasive Technology
(Philadelphia, PA)
|
Family
ID: |
46255782 |
Appl.
No.: |
08/799,206 |
Filed: |
February 13, 1997 |
Related U.S. Patent Documents
|
|
|
|
|
|
|
Application
Number |
Filing Date |
Patent Number |
Issue Date |
|
|
731443 |
Oct 15, 1996 |
|
|
|
|
031945 |
Mar 16, 1993 |
5564417 |
|
|
|
076370 |
Jun 14, 1993 |
5553614 |
|
|
|
645590 |
Jan 24, 1991 |
|
|
|
|
Current U.S.
Class: |
600/310; 600/323;
600/336 |
Current CPC
Class: |
A61B
5/0075 (20130101); A61B 5/0084 (20130101); A61B
5/14551 (20130101); A61B 5/14552 (20130101); A61B
5/14553 (20130101); A61B 5/1459 (20130101); A61B
5/6828 (20130101); A61B 5/7228 (20130101); G01N
21/3151 (20130101); G01N 21/4795 (20130101); G01N
21/49 (20130101); A61B 2562/0233 (20130101); A61B
2562/0242 (20130101); A61B 2562/043 (20130101); G01N
21/47 (20130101); G01N 2021/1789 (20130101); G01N
2021/3181 (20130101); G01N 2201/0696 (20130101) |
Current International
Class: |
A61B
5/00 (20060101); G01N 21/64 (20060101); G01N
21/47 (20060101); G01N 21/31 (20060101); G01N
21/49 (20060101); A61B 005/00 () |
Field of
Search: |
;600/310,604,322-343,407,473,476,472 |
References Cited
[Referenced By]
U.S. Patent Documents
Primary Examiner: Shay; David M.
Attorney, Agent or Firm: Fish & Richardson, P.C.
Parent Case Text
CROSS REFERENCE TO RELATED APPLICATION
This application is a continuation-in-part of U.S. patent
application Ser. No. 08/731,443, filed Oct. 15, 1996; which in turn
is a continuation of U.S. patent application Ser. No. 08/031,945,
filed Mar. 16, 1993, now U.S. Pat. No. 5,564,417; which in turn is
a continuation-in-part of U.S. patent application Ser. No.
08/076,370, filed Jun. 14, 1993, issued as U.S. Pat. No. 5,553,614;
which is a continuation of U.S. patent application Ser. No.
07/645,590, filed Jan. 24, 1991, now abandoned, all of which are
incorporated by reference as if fully set forth herein.
Claims
What is claimed is:
1. An in vivo spectroscopic method for examining biological tissue,
comprising:
positioning an input port at a selected location relative to the
biological tissue;
positioning a detection port at another location spaced at a
selected distance of several centimeters from said input port;
generating a first carrier waveform at a selected frequency on the
order of 10.sup.8 Hz;
introducing into the tissue at said input port electromagnetic
radiation of a selected wavelength modulated by said carrier
waveform, said wavelength being sensitive to concentration of said
absorptive pigment present in the tissue;
detecting at said detection port the radiation that has migrated
over migration paths in a portion of the tissue from said input
port, said portion of the tissue depending on said locations of
said input and detection ports;
comparing the detected radiation with the introduced radiation and
measuring therefrom a phase shift (.theta.) of said detected
radiation;
providing a scattering of said portion of the tissue;
using a recursive procedure to calculate a value of the absorption
coefficient, at said wavelength, using the following equation:
##EQU14##
and
examining said tissue portion by using said calculated value of the
absorption coefficient.
2. The spectroscopic method of claim 1 wherein said step of
providing said scattering property includes
generating a second carrier waveform at a second selected frequency
on the order of 10.sup.8 Hz; and
measuring said phase shift at said second frequency for each said
wavelength.
3. The spectroscopic method of claim 1 wherein said absorptive
pigment is oxyhemoglobin or deoxyhemoglobin.
4. The method of claim 1 wherein said examining includes
quantifying a concentration of said absorptive pigment in said
tissue portion by using said calculated value of the absorption
coefficient.
5. The spectroscopic method of claim 1 further comprising;
introducing into the tissue at said input port electromagnetic
radiation of a second selected wavelength modulated by said carrier
waveform, at least one of said wavelengths being sensitive to
concentration of said absorptive pigment present in the tissue,
said tissue exhibiting a similar scattering property at said
wavelengths;
detecting at said detection port the radiation of said second
wavelength that has migrated over migration paths in a portion of
the tissue from said input port;
comparing, at said second wavelength, the detected radiation with
the introduced radiation and measuring therefrom a phase shift
(.theta.) of said detected radiation at said second wavelength;
using the recursive procedure to calculate a value of the
absorption coefficient, at said second wavelength, using the
following equation: ##EQU15##
and
said examining includes quantifying the concentration of said
absorptive pigment in said tissue portion by using said calculated
values of the absorption coefficients at said two wavelengths.
6. The spectroscopic method of claim 5 wherein said quantifying
step includes
calculating a ratio of absorption coefficients at said two
wavelengths; and
calculating a value of oxygen saturation based on said ratio.
7. The spectroscopic method of claim 1 wherein said step of
providing said scattering property includes looking up a value of
said scattering property from a lookup table that includes said
values for different tissue types.
8. The spectroscopic method of claim 7 wherein said value of said
scattering property is the effective scattering coefficient
(1-g).mu..sub.s.
9. The method of claim 1 further including detecting the radiation
that has migrated over migration paths in another said portion of
the tissue and performing said comparing, said providing and said
using the recursive procedure to calculate a value of the
absorption coefficient at said wavelength.
10. The method of claim 9 repeated for several said portions of the
tissue.
11. A spectroscopic system for examining biological tissue,
comprising:
an oscillator constructed to generate a first carrier waveform of a
first frequency on the order of 10.sup.8 Hz;
a light source, operatively coupled to said oscillator, constructed
to generate electromagnetic radiation of a selected wavelength
modulated by said carrier waveform, said wavelength being sensitive
to concentration of said absorptive pigment present in the
tissue;
an input port constructed to introduce said radiation into the
tissue;
a detection port, located several centimeters apart from said input
port, constructed to acquire photons of the radiation that has
migrated from said input port over migration paths in a portion of
the tissue, said portion of the tissue depending on locations of
said input and detection ports;
a detector, optically connected to said detection port, constructed
to detect the radiation;
a phase detector constructed to compare the detected radiation with
the introduced radiation and determine therefrom the phase shift of
said detected radiation; and
a processor constructed to receive said phase shift and a
scattering property of said portion of the tissue, processor being
arranged to use a recursive procedure to calculate a value of the
absorption coefficient, at said wavelength, using the following
equation: ##EQU16##
12. The system of claim 11 including
a second oscillator constructed to generate a second carrier
waveform of a second selected frequency on the order of 10.sup.8
Hz;
said source operatively coupled to said second oscillator,
constructed to generate electromagnetic radiation of said
wavelength modulated by said second carrier waveform;
said detector further constructed to detect the radiation modulated
by said second carrier waveform; and
said phase detector further constructed to compare, at said
wavelength, the detected radiation of said second carrier waveform
with the introduced radiation and determine therefrom the phase
shift of said detected radiation.
13. The system of claim 11 wherein said absorptive pigment is
oxyhemoglobin or deoxyhemoglobin.
14. The system of claim 11 wherein said processor is further
arranged to quantify a concentration of said absorptive pigment in
said tissue portion by using said calculated value of the
absorption coefficient.
15. The system of claim 11 wherein said processor is arranged to
use said recursive procedure that involves the Newton-Raphson
method.
16. The system of claim 13 wherein said light source is constructed
to generate electromagnetic radiation of a second wavelength
modulated by said carrier waveform, at least one of said
wavelengths being sensitive to said concentration of said
absorptive pigment present in the tissue, said tissue exhibiting a
similar scattering property at said wavelengths; said detector
being constructed to detect the radiation at said second
wavelength; said phase detector being constructed to compare the
detected radiation with the introduced radiation and determine
therefrom the phase shift at said second wavelength; and said
processor being constructed to receive said phase shift at said
second wavelength and use the recursive procedure to calculate a
value of the absorption coefficient, at said second wavelength,
using the following equation: ##EQU17##
17. The system of claim 16 wherein said processor is constructed to
calculate a ratio of absorption coefficients at said two
wavelengths determined by the recursive procedure, and calculate a
value of oxygen saturation based on said ratio.
18. The system of claim 11 further including a look up table
comprising values of said scattering property for different tissue
types.
19. The system of claim 18 wherein said value of said scattering
property is the effective scattering coefficient, (1-g)
.mu..sub.s.
20. An in vivo spectroscopic method for examining biological
tissue, comprising:
positioning an input port at a selected location relative to the
biological tissue;
positioning a detection port at another location spaced at a
selected distance of several centimeters from said input port;
generating a first carrier waveform at a selected frequency on the
order of 10.sup.8 Hz;
introducing into the tissue at said input port electromagnetic
radiation of a selected wavelength modulated by said carrier
waveform, said wavelength being sensitive to concentration of said
absorptive pigment present in the tissue;
detecting at said detection port the radiation that has migrated
over migration paths in a portion of the tissue from said input
port, said portion of the tissue depending on said locations of
said input and detection ports;
creating a first and a second reference phase signals of predefined
substantially different phases;
comparing the detected radiation with said first and said second
reference signals and determining therefrom a real output signal
and an imaginary output signal, respectively;
providing said scattering property of said portion of the
tissue;
calculating a value of the phase shift (.theta.) of said detected
radiation as the inverse tangent of the ratio of said imaginary
output signal and said real output signal;
using a recursive procedure to calculate a value of the absorption
coefficient, at said wavelength, using the following equation:
##EQU18##
and
examining said tissue portion by using said calculated value of the
absorption coefficient.
21. The method of claim 20 wherein said examining includes
quantifying a concentration of said absorptive pigment in said
tissue portion by using said calculated value of the absorption
coefficient.
22. The method of claim 20 further including detecting the
radiation that has migrated over migration paths in another said
portion of the tissue and performing said comparing, said
providing, said calculating and said using the recursive procedure
to calculate a value of the absorption coefficient at said
wavelength.
23. The method of claim 22 repeated for several said portions of
the tissue.
24. The spectroscopic method of claim 20 further comprising
introducing into the tissue at said input port electromagnetic
radiation of a second selected wavelength modulated by said carrier
waveform, at least one of said wavelengths being sensitive to
concentration of said absorptive pigment present in the tissue,
said tissue exhibiting a similar scattering property at said
wavelengths;
detecting at said detection port the radiation of said second
wavelength that has migrated over migration paths in a portion of
the tissue from said input port;
comparing the detected radiation of said second wavelength with
said first and said second reference signals and determining
therefrom a real output signal and an imaginary output signal,
respectively, at said second wavelength;
calculating a value of the phase shift (.theta.) of said detected
radiation at said second wavelength as the inverse tangent of the
ratio of said imaginary output signal and said real output
signal
using the recursive procedure to calculate a value of the
absorption coefficient, at said second wavelength, using the
following equation: ##EQU19##
and
said examining includes quantifying concentration of said
absorptive pigment in said tissue portion by using said calculated
values of the absorption coefficients at said two wavelengths.
25. The spectroscopic method of claim 24 wherein said quantifying
step includes
calculating a ratio of absorption coefficients at said two
wavelengths; and
calculating a value of oxygen saturation based on said ratio.
26. The spectroscopic method of claim 24 wherein said quantifying
step includes calculating, at each wavelength, a detected amplitude
(A) as a square root of a sum of squares of said real output signal
and said imaginary output signal.
27. The spectroscopic method of claim 24 wherein said step of
providing said scattering property includes
generating a second carrier waveform at a second selected frequency
on the order of 10.sup.8 Hz; and
calculating (.theta..sub..lambda.) said phase shift at said second
frequency for each said wavelength.
28. A spectroscopic system for examining biological tissue,
comprising:
an oscillator constructed to generate a first carrier waveform at a
selected frequency on the order of 10.sup.8 Hz;
a light source, operatively coupled to said first oscillator,
constructed to generate electromagnetic radiation of a selected
wavelength modulated by said carrier waveform, said wavelength
being sensitive to concentration of said absorptive pigment present
in the tissue;
an input port constructed to introduce photons of electromagnetic
radiation into the examined biological tissue;
a detection port, spaced several centimeters apart from said input
port, constructed to acquire photons that have migrated over
migration paths in an examined portion of the tissue from said
input port, said portion of the tissue depending on locations of
said input and detection ports;
a detector constructed to detect, at said detection port, the
radiation that has migrated over migration paths in the examined
portion of the tissue;
a phase splitter constructed to receive said carrier waveform and
produce first and second reference phase signals of predefined
substantially different phases;
first and second double balanced mixers connected to receive from
said phase splitter said first and second reference phase signals,
respectively, and connected to receive from said detector said
detector signal, and constructed to produce therefrom a real output
signal and an imaginary output signal, respectively; and
a processor constructed to receive a scattering property of said
portion of the tissue and arranged to calculate a phase shift
(.theta.) of said detected radiation as the inverse tangent of the
ratio of said imaginary output signal and said real output signal,
said processor being further arranged to use a recursive procedure
to calculate a value of the absorption coefficient, at said
wavelength, using the following equation: ##EQU20##
29. The system of claim 28 including
a second oscillator constructed to generate a second carrier
waveform of a second selected frequency on the order of 10.sup.8
Hz;
said source operatively coupled to said second oscillator,
constructed to generate electromagnetic radiation of said
wavelength modulated by said second carrier waveform;
said detector further constructed to detect the radiation modulated
by said second carrier waveform; and
said first and second double balanced mixers connected to receive
from said phase splitter said first and second reference phase
signals, respectively, and connected to receive from said detector
said detector signal, and constructed to produce therefrom a real
output signal and an imaginary output signal, respectively, at said
second frequency.
30. The system of claim 28 wherein said processor is further
arranged to quantify a concentration of said absorptive pigment in
said tissue portion by using said calculated value of the
absorption coefficient.
31. The system of claim 28 wherein said absorptive pigment is
oxyhemoglobin or deoxyhemoglobin.
32. The system of claim 28 further including a look up table
comprising values of said scattering property for different tissue
types.
33. The system of claim 32 wherein said value of said scattering
property is the effective scattering coefficient, (1-g)
.mu..sub.s.
34. The system of claim 28 wherein said light source is constructed
to generate electromagnetic radiation of a second wavelengths
modulated by said carrier waveform, at least one of said
wavelengths being sensitive to said concentration of said
absorptive pigment present in the tissue, said tissue exhibiting a
similar scattering property at said wavelengths; said detector
being constructed to detect the radiation at said second
wavelength; said first and second double balanced mixers being
connected to receive from said phase splitter said first and second
reference phase signals, respectively, and connected to receive
from said detector said detector signal, and constructed to produce
therefrom, at said second wavelength, a real output signal and an
imaginary output signal, respectively; and said processor being
arranged to calculate a phase shift (.theta.) of said detected
radiation at said second wavelength, as the inverse tangent of the
ratio of said imaginary output signal and said real output signal,
said processor being also arranged to use the recursive procedure
to calculate a value of the absorption coefficient, at said second
wavelength, using the following equation: ##EQU21##
35. The system of claim 34 wherein one of said wavelengths is
sensitive to oxygenation of hemoglobin and said processor is
constructed to calculate a ratio of said absorption coefficients,
at said two wavelengths, and calculate therefrom a value of oxygen
saturation based on said ratio.
36. The system of claim 34 wherein said processor is arranged to
use said recursive procedure that involves the Newton-Raphson
method.
Description
BACKGROUND OF THE INVENTION
The present invention relates to quantitative analyses of
absorptive constituents in biological tissues by employing a phase
modulation spectroscopy.
Continuous wave (CW) tissue oximeters have been widely used to
determine in vivo concentration of an optically absorbing pigment
(e.g., hemoglobin, oxyhemoglobin) in biological tissue. The CW
oximeters measure attenuation of continuous light in the tissue and
evaluate the concentration based on the Beer Lambert equation or a
modified Beer Lambert absorbance equation. The Beer Lambert
equation (1) describes the relationship between the concentration
of an absorbent constituent (C), the extinction coefficient
(.epsilon.), the photon migration pathlength <L>, and the
attenuated light intensity (I/I.sub.o). ##EQU1##
The CW spectrophotometric techniques can not determine .epsilon.,
C, and <L> at the same time. If one could assume that the
photon pathlength were constant and uniform throughout all
subjects, direct quantitation of the constituent concentration (C)
using CW oximeters would be possible.
In tissue, the optical migration pathlength varies with the size,
structure, and physiology of the internal tissue examined by the CW
oximeters. For example, in the brain, the gray and white matter and
the structures thereof are different in various individuals. In
addition, the photon migration pathlength itself is a function of
the relative concentration of absorbing constituents. As a result,
the pathlength through an organ with a high blood hemoglobin
concentration, for example, will be different from the same with a
low blood hemoglobin concentration. Furthermore, the pathlength is
frequently dependent upon the wavelength of the light since the
absorption coefficient of many tissue constituents is wavelength
dependent. Thus, where possible, it is advantageous to measure the
pathlength directly when quantifying the hemoglobin concentration
in tissue.
SUMMARY OF THE INVENTION
In general, in one aspect, a spectroscopic system for quantifying
in vivo concentration of an absorptive pigment in biological tissue
includes an oscillator constructed to generate a first carrier
waveform of a first frequency on the order of 10.sup.8 Hz (i.e., in
the range of 10 MHz to 1 GHz), a light source, operatively coupled
to the oscillator, constructed to generate electromagnetic
radiation of a selected wavelengths modulated by the carrier
waveform, and a detector constructed to detect radiation that has
migrated over photon migration paths in the tissue from an input
port to a detection port spaced several centimeters apart. The
wavelength is sensitive to concentration of the absorptive pigment
present in the tissue. A phase detector is constructed to compare
the detected radiation with the introduced radiation and determine
therefrom the phase shift of the detected radiation at each
wavelength. A processor is constructed to receive the phase shift
and a scattering property of the portion of the tissue and
calculate a value of the absorption coefficient, at the wavelength,
using Eq. 4.
In another embodiment, the spectroscopic system includes a light
source further constructed to generate electromagnetic radiation of
a second wavelengths modulated by the carrier waveform. At least
one of the wavelengths is sensitive to concentration of an
absorptive pigment present in the tissue, while the tissue exhibits
similar scattering properties at the two wavelengths. The detector
is constructed to detect the radiation at the second wavelength.
The phase detector is constructed to compare the detected radiation
with the introduced radiation and determine therefrom the phase
shift at the second wavelength. The processor is constructed to
receive the phase shift at the second wavelength and calculate a
value of the absorption coefficient, at the second wavelength,
using Eq. 4.
In general, in one aspect, a spectroscopic system for quantifying
in vivo concentration of an absorptive pigment in biological tissue
includes an oscillator constructed to generate a first carrier
waveform of a first frequency on the order of 10.sup.8 Hz (i.e., in
the range of 10 MHz to 1 GHz), a light source, operatively coupled
to the oscillator, constructed to generate electromagnetic
radiation of a selected wavelengths modulated by the carrier
waveform, and a detector constructed to detect radiation that has
migrated over photon migration paths in the tissue from an input
port to a detection port spaced several centimeters apart. The
wavelength is sensitive to concentration of the absorptive pigment
present in the tissue. The spectroscopic system also includes a
phase splitter, two double balanced mixers, and a processor. The
phase splitter is constructed to receive the carrier waveform and
produce first and second reference phase signals of predefined
substantially different phases. The first and second double
balanced mixers are connected to receive from the phase splitter
the first and second reference phase signals, respectively, and are
connected to receive from the detector the detector signal, in
order to produce therefrom a real output signal and an imaginary
output signal, respectively. The processor is constructed to
receive a scattering property of the portion of the tissue and is
constructed to quantify the concentration of the absorptive pigment
by calculating phase shift (.theta.) of the detected radiation as
the inverse tangent of the ratio of the imaginary output signal and
the real output signal. The processor also calculates a value of
the absorption coefficient, at the wavelength, using Eq. 4.
In another embodiment, the spectroscopic system includes a light
source further constructed to generate electromagnetic radiation of
a second wavelengths modulated by the carrier waveform. At least
one of the wavelengths is sensitive to concentration of an
absorptive pigment present in the tissue, while the tissue exhibits
similar scattering properties at the two wavelengths. The detector
is constructed to detect the radiation at the second wavelength.
The first and second double balanced mixers are connected to
receive from the phase splitter the first and second reference
phase signals, respectively, and connected to receive from the
detector the detector signal at the second wavelength. The mixers
are constructed to produce therefrom a real output signal and an
imaginary output signal, respectively, at the second wavelength.
The processor is constructed to quantify the concentration of the
absorptive pigment by calculating phase shift (.theta.) of the
detected radiation as the inverse tangent of the ratio of the
imaginary output signal and the real output signal and by
calculating a value of the absorption coefficient, at the
wavelength, using Eq. 4.
As different embodiments, the spectrophotometer may be a dual
wavelength, single frequency system or a dual wavelength, dual
frequency system. Each system can measure data for a single
source-detector separation (i.e., separation of the input port and
the detection port) or for several source-detector separations.
Different embodiments of the spectrophotometer may include one or
more of the following features.
The spectrophotometer may include a second oscillator constructed
to generate a second carrier waveform of a second selected
frequency on the order of 10.sup.8 Hz, while the tissue exhibits
similar scattering properties at the selected frequencies. The
source is operatively coupled to the second oscillator and is
constructed to generate electromagnetic radiation of the two
wavelengths modulated by the second carrier waveform. The detector
is further constructed to detect the radiation modulated by the
second carrier waveform. The phase detector is further constructed
to compare, at each the wavelength, the detected radiation of the
second carrier waveform with the introduced radiation and determine
therefrom the phase shift of the detected radiation.
The processor may calculate a ratio of absorption coefficients at
the two wavelengths, and calculate a value of oxygen saturation
based on the ratio.
The spectrophotometer may include a look up table including values
of the scattering property for different tissue types. These values
may be the effective scattering coefficient, (1-g).mu..sub.s.
The absorptive pigment may be oxyhemoglobin or deoxyhemoglobin.
BRIEF DESCRIPTION OF THE DRAWING
FIG. 1 is a block diagram of a pathlength corrected oximeter in
accordance with the present invention.
FIG. 2 is a schematic circuit diagram of a 50.1 MHz (50.125 MHz)
oscillator used in the oximeter of FIG. 1.
FIG. 3 is a schematic circuit diagram of a PIN diode and a
preamplifier used in the oximeter of FIG. 1.
FIG. 4 is a schematic circuit diagram of a magnitude detector used
in the oximeter of FIG. 1.
FIG. 5 is a schematic circuit diagram of a 25 kHz filter used in
the oximeter of FIG. 1.
FIG. 6 is a schematic diagram of an AGC circuit of the oximeter of
FIG. 1.
FIG. 7 is a schematic circuit diagram of a phase detector of the
oximeter of FIG. 1.
FIG. 8A is a plan view of a source-detector probe of the
oximeter.
FIG. 8B is a transverse cross-sectional view taken on lines 8B of
FIG. 8A further showing the photon migration.
FIG. 9 is a block diagram of another embodiment of a phase
modulation spectrophotometer.
FIGS. 10A and 10B display simulation results for oxygen saturation
values and their noise dependence, respectively, calculated by
using a high frequency approximation.
FIGS. 11A and 11B display simulation results for oxygen saturation
values and their noise dependence, respectively, calculated by
using a low frequency approximation.
FIG. 12 displays simulation results for oxygen saturation values as
a function of a varying scattering coefficient.
FIGS. 13A and 13B display raw data and calculated saturation data,
respectively, measured on a newborn piglet.
DESCRIPTION OF THE PREFERRED EMBODIMENTS
One preferred embodiment of the pathlength corrected oximeter
utilizes three LEDs for generation of light at three selected
wavelengths intensity modulated at a frequency of 50.1 MHz and
coupled directly to the examined tissue. At each wavelength, the
introduced light is altered by the tissue and is detected by a wide
area photodiode placed against the skin. The introduced and
detected radiations are compared to determine their relative phase
shift that corresponds to an average pathlength of the migrating
photons and, furthermore, the light attenuation is determined.
Referring to FIG. 1, the oximeter includes a master oscillator 10
operating at 50.1 MHz connected to a power amplifier 15 of
sufficient output power to drive LEDs 22a, 22b, and 22c (for
example HLP 20RG or HLP 40RG made by Hitachi) that emit 760 nm, 840
nm, and 905 nm (or 950 nm) light, respectively. A second local
oscillator 14 operating at 50.125 MHz and mixer 12 are used to
generate a reference frequency 13 of 25 kHz. Each LED directly
positioned on the skin has an appropriate heat sink to eliminate
uncomfortable temperature increases that could also alter blood
perfusion of the surrounding tissue. Three PIN diode detectors 24a,
24b, and 24c are placed at a distance of approximately 5 cm from
the LEDs and have a detection area of about 1 cm.sup.2. Photons
migrating a few centimeters deep into the tissue are detected by
the respective PIN diodes. The source-detector separation can be
increased or decreased to capture deeper or shallower migrating
photons. The signals from PIN diodes 24a, 24b, and 24c are
amplified by preamplifiers 30a, 30b, and 30c, respectively.
The amplified signals (32a, 32b, 32c) are sent to magnitude
detectors 36a, 36b, and 36c and to mixers 40a, 40b, and 40c,
respectively. The magnitude detectors are used to determine
intensity values of detected signals at each wavelength to be used
in Eq. 1. Each mixer, connected to receive a 50.125 MHz reference
signal (41a, 41b, 41c) from local oscillator 14, converts the
detection signal to a 25 kHz frequency signal (42a, 42b, 42c). The
mixers are high dynamic range frequency mixers, model SRA-1H,
commercially available from Mini-Circuits (Brooklyn N.Y.). The
detection signals (42a, 42b, and 42c) are filtered by filters 45a,
45b, 45c, respectively.
Phase detectors 60a, 60b, and 60c are used to determine phase shift
between the input signal and the detected signal at each
wavelength. Each phase detector receives the 25 kHz detection
signal (54a, 54b, 54c) and the 25 kHz reference signal (56a, 56b,
56c), both of which are automatically leveled by automatic gain
controls 50 and 52 to cover the dynamic range of signal changes.
Phase detectors 60a, 60b, and 60c generate phase shift signals
(62a, 62b, 62c) corresponding to the migration delay of photons at
each wavelength. Each phase shift signal is proportional to the
migration pathlength used in calculation algorithms performed by
processor 70.
FIG. 2 shows a schematic circuit diagram of a precision oscillator
used as the 50.1 MHz master oscillator 10 and 50.125 MHz local
oscillator 14. The oscillator crystals are neutralized for
operation in the fundamental resonance mode; this achieves
long-term stability. Both oscillators are thermally coupled so that
their frequency difference is maintained constant at 25 kHz if a
frequency drift occurs.
PIN diodes 24a, 24b, and 24c are directly connected to their
respective preamplifiers 30a, 30b, and 30c, as shown in FIG. 3. The
oximeter uses PIN silicon photodiodes S1723-04 with 10 mm.times.10
mm sensitive area and spectral response in the range of 320 nm to
1060 nm. The detection signal is amplified by stages 29 and 31,
each providing about 20 dB amplification. The NE5205N operational
amplifier is powered at +8V to operate in a high gain regime. The
8V signal is supplied by a voltage regulator 33. The amplified
detection signals (32a, 32b, and 32c) are sent to magnitude
detectors 36a, 36b, and 36c, shown in FIG. 4. The magnitude values
(37a, 37b, and 37c) are sent to processor 70 that calculates the
light attenuation ratio or logarithm thereof as shown Eq. 1.
Also referring to FIG. 5, the AGC circuit uses MC 1350 integrated
circuit for amplification that maintains the input signal of phase
detector 60 at substantially constant levels. The amount of gain is
selected to be equal for AGCs, 50 and 52. The signal amplitude is
controlled by a feedback network 53. The AGCs provide a
substantially constant amplitude of the detected and reference
signals to eliminate variations in the detected phase shift due to
cross talk between amplitude and phase changes in the phase
detector.
Referring to FIG. 6, each phase detector includes a Schmitt trigger
that converts the substantially sinusoidal detection signal (54a,
54b, 54c) and reference signal (56a, 56b, 56c) to square waves. The
square waves are input to a detector that has complementary MOS
silicon-gate transistors. The phase shift signal is sent to
processor 70.
The oximeter is calibrated by measuring the phase shift for a
selected distance in a known medium, i.e., using a standard delay
unit, and by switching the length of a connector wire to change the
electrical delay between master oscillator 10 and local oscillator
14.
Referring to FIGS. 8A and 8B source-detector probe 20 includes
several LEDs (22a, 22b, 22c) of selected wavelengths and PIN
photodiodes (24a, 24b, 24c) mounted in a body-conformable support
structure 21. Structure 21 also includes a photon escape barrier 27
made of a material with selected scattering and absorption
properties (for example, styrofoam) designed to return escaping
photons back to the examined tissue. The support structure further
includes a second conformable barrier 28, located between the LEDs
and the diode detectors, designed to absorb photons directly
propagating from the source to the detector and thus prevent
detection of photons that migrate subcutaneously. Support structure
21 also includes electronic circuitry 29 encapsulated by an
electronic shield 21a.
Each PIN diode is provided with an evaporated single wavelength
film filter (25a, 25b, 25c). The filters eliminate the cross talk
of different wavelength signals and allow continuous operation of
the three light sources, i.e., no time sharing is needed.
The use of photodiode detectors has substantial advantages when
compared with the photomultiplier tube used in standard phase
modulation systems. The photodiodes are placed directly on the
skin, i.e., no optical fibers are needed. Furthermore, there is no
need to use a high voltage power supply that is necessary for the
photomultiplier tube. The photodiodes are much smaller and are easy
to place close to the skin. Advantages of the photomultiplier tube
are a huge multiplication gain and a possibility of direct mixing
at the photomultiplier; this cannot be achieved directly by a
photodiode. This invention envisions the use of several different
photodiodes such as PIN diode, avalanche diode, and other.
The processor uses algorithms that are based on equations described
by E. M. Sevick et al. in "Quantitation of Time- and
Frequency-Resolved Optical Spectra for the Determination of Tissue
Oxygenation," published in Analytical Biochemistry 195, 330, Apr.
15, 1991, which is incorporated by reference as if fully set forth
herein. The photon migration in biological tissue is a diffusional
process in which the photon fluence rate, .phi.(r,t), is
distributed from the source. The fluence rate is equal to
N.sub..alpha. c, or the product of the number of the photon at
position r and time, t, and the speed of photons through the
medium. The fluence rate, or the effective "concentration" of
photons at position r and time t, in the tissue or turbid media may
be obtained from the solution of the general diffusion equation
##EQU2##
where D is the diffusion coefficient and S a source term. For
photon migration, the diffusion coefficient is equal to
##EQU3##
where .mu..sub.s is the scattering coefficient (cm.sup.-1) and g is
the mean cosine of scattering angle. The term (1-g).mu..sub.s is
referred to as the effective scattering coefficient and is equal to
the reciprocal of the isotropic, mean scattering length, l* (i.e.,
when the direction of scatter is completely random). The absorption
coefficient .mu..sub.a is based upon the Napierian extinction
coefficient.
The source at .rho.=0 consists of light whose intensity is
sinusoidally modulated at a frequency f. The light intensity
detected at a distance .rho. away from the source is both amplitude
demodulated and phase shifted with respect to the incident source
intensity. The measured phase shift, .theta., and the modulation,
M, of the detected light with respect to that of the incident light
characterize the tissue wherein the detected photons migrated over
a distribution of pathlengths. The phase shift describes the
pathlength distribution in the frequency domain. It can be directly
related to the mean of the distribution of pathlengths traveled by
photons before detection. The modulation of the detected intensity
also varies with changes in the absorbance and pathlength
distribution. Pathlengths can be used to detect changes in
absorption in strongly scattering media. Modulation may also be
used to detect changes in absorption in the tissue. In phase
modulation (frequency modulation), the source term represents a
sinusoidally modulated photon flux at point
.rho.=0;S(.rho.=0,t)=A+M.multidot.sin (2.pi.f.multidot.t).
Expressions of the phase shift and modulation of the detected
intensity may also be directly found from Eq. 2.
The analytical solution for .theta. and M can be obtained from the
sine and cosine Fourier transforms of Eq. 2: ##EQU4##
At each wavelength, for low modulation frequencies, i.e.,
2.pi.f<<.mu..sub.a.multidot.c, the phase shift
(.theta..sup..lambda.) (62a, 62b, 62c) is used to calculate the
pathlength as follows: ##EQU5##
wherein f is modulation frequency of the introduced light which is
in the range of 10 MHz to 100 MHz; t.sup..lambda. is the photon
migration delay time; c is the speed of photons in the scattering
medium; and L.sup..lambda. is the migration pathlength. The
modulation frequency of 50 MHz was selected due to the frequency
limitation of the LEDs and photodiodes. However, for faster LEDs
and photodiodes it may be desirable to use higher modulation
frequencies that increase the phase shift resolution.
At high modulation frequencies, i.e.,
2.pi.f>>.mu..sub.a.multidot.c, the phase shift is no longer
proportional to the mean time of flight <t>. ##EQU6##
wherein .rho. is the source-detector separation;
a=(6.pi./c).sup.1/2 sin.pi./4; (1-g) .mu..sub.s is the effective
scattering coefficient, .mu..sub.a.sup..lambda. is the absorption
coefficient at wavelength .lambda., .alpha..sup..lambda. is the
background absorbance at wavelength .lambda., and
.theta..sub.0.sup..lambda. thus represents background scattering
and absorption. At two wavelengths, the ratio of absorption
coefficients is determined as follows: ##EQU7##
The wavelengths are in the visible and infra-red range and are
selected to have absorbance sensitive (or insensitive) to various
tissue components such as water, cytochrome iron and copper, oxy-
and deoxygenated forms of hemoglobin, myoglobin, melanin, glucose
and other.
For oxygenated and deoxygenated hemoglobin, the absorption
coefficient written in terms of Beer Lambert relationship is as
follows:
.mu..sub.a.sup..lambda..sup..sub.1
=.epsilon..sub.Hb.sup..lambda..sup..sub.1
[Hb]+.epsilon..sub.HbO.sup..lambda..sup..sub.1 [HbO.sub.2
]+.alpha..sup..lambda..sup..sub.1 (12)
wherein .epsilon..sub.Hb.sup..lambda.1 and
.epsilon..sub.HbO.sup..lambda.1. are extinction coefficients for
hemoglobin and deoxyhemoglobin that can be stored in a look up
table; [Hb], [Hb0.sub.2 ] are the tissue concentration of
hemoglobin and oxyhemoglobin, respectively; .alpha..sup..lambda.1
is background absorbance at wavelength .lambda..sub.1.
Tissue hemoglobin saturation can also be determined from
dual-wavelength, dual-frequency measurements of the phase shift.
For high modulation frequencies, (2.pi.f.sub.1
>>.mu..sub.a.sup..lambda.1 c and f.sub.2
>>.mu..sub.a.sup..lambda.2 c) the differences in the measured
phase shift at one wavelength and two frequencies can be expressed
as ##EQU8##
The ratio of this difference measured at two wavelengths can thus
be written ##EQU9##
Since the scattering coefficient is wavelength-insensitive over the
near-infrared range employed, this dual-frequency, dual-wavelength
phase modulated spectroscopy can be used to obtain the ratio of
absorption coefficients.
Furthermore, as predicted from the diffusion approximation, the
magnitude of the phase shift increases with the source-detector
separation, .rho.. Thus, in homogeneous tissues, the phase shifts
measured for several .rho. can be used to calculate the absorption
and scattering coefficients. These coefficients can be used either
by employing Eq. 4 or the equations for the high and low
approximations. Similarly, the magnitude of the detected radiation
can be measured for different source-detector separations, and the
absorption and scattering coefficients can be calculated by using
Eq. 5.
The hemoglobin saturation is conventionally defined as follows:
##EQU10##
For a three wavelength measurement, the hemoglobin saturation can
be calculated using Eqs. (12) and (15) as follows: ##EQU11##
Thus, processor 70 determines Y from the above equations for each
wavelength .lambda..sub.1, .lambda..sub.2, .lambda..sub.3.
In another embodiment, the spectrophotometer's electronics includes
a low frequency module suitably and a high frequency module
switchably coupled to the same source-detector probe 20. The low
frequency module and the arrangement of the source-detector probe
are substantially similar to the hemoglobinometer described in a
co-pending U.S. patent application Ser. No. 701,127 filed May 16,
1991 which is incorporated by reference as if fully set forth
herein. The low frequency module corresponds to a standard oximeter
with modulation frequencies in the range of a few hertz to 10.sup.4
hertz and is adapted to provide intensity attenuation data at two
or three wavelengths. Then, the LEDs are switched to the high
frequency phase modulation unit, similar to the unit of FIG. 1,
which determines the average pathlength at each wavelength. The
attenuation and pathlength data are sent to processor 70 for
determination of a physiological property of the examined
tissue.
In another embodiment, the pathlength corrected oximeter utilizes
the same LED sources (22a, 22b, 22c) sinusoidally modulated at a
selected frequency comparable to the average migration time of
photons scattered in the examined tissue on paths from the optical
input port of the LED's to the optical detection part of the
photodiode detectors (24a, 24b, 24c), but the electronic circuitry
is different. Referring to FIG. 9, this embodiment utilizes a 200
MHz precision oscillator 61, which drives two laser diodes 62 and
64, again at 760 and 816 nm. The outputs of the laser diodes are
time shared into filter optic coupling 68 and the head 70. Detector
72 provides output to an amplifier 74 and to two wide band double
balance mixers (DBM) 76 and 78 which are coupled through a
90.degree. phase splitter 80 so that real (R) and imaginary (I)
portions of the signal are obtained. The double balance mixers 76
and 78 preferably operate at the modulation frequency. The phase
(.theta..sup..lambda.) is the angle whose tangent is the imaginary
over the real part. ##EQU12##
The amplitude is the square root of the sum of the squares of these
values, providing the phase shift has been taken out as the
residual phase shift .theta. set to zero.
This embodiment uses summing and dividing circuits to calculate the
modulation index, which is the quotient of the amplitude over the
amplitude plus the DC component obtained from a narrow band
detector 82. ##EQU13##
The phase processor receives the phase shifts for the phase and
amplitude values for two or three wavelengths and calculates the
ratio of the phase shifts. For each wavelength, the phase shift and
the DC amplitude are used to determine a selected tissue property,
e.g., hemoglobin oxygenation.
To study the influence of variation in the scattering coefficient
on the quantitation of the absorption measurements, several
simulations were performed. The simulations assumed the phase shift
measurements at two wavelengths and several frequencies (10 MHz, 50
MHz, 200 MHz and 500 MHz). Hemoglobin saturation levels (Y) were
varied in the range of 5%.ltoreq.Y.ltoreq.100%, and the absorption
coefficients were varied in the range of
0.5.ltoreq..mu..sub.a.ltoreq.1.5 cm.sup.-1, while the scattering
coefficient .mu..sub.s '=5 cm.sup.-1 was kept constant; these
values correspond to typical values for human tissue. FIGS. 10A and
10B show simulation results obtained by using the high frequency
approximation (2.pi.f>>.mu..sub.a c) for modulation
frequencies f=50, 200 and 500 MHz, assuming
.theta..sub.0.sup..lambda.1 =.theta..sub.0.sup..lambda.2
=.theta..sub.0, and .mu..sub.a
c.apprxeq.2.multidot.10.sup.9.multidot..theta..sub.0. As shown in
FIG. 10A, the calculated saturation error decreases with frequency,
but still introduces a significant error even for the 500 MHz at
low saturation values. FIG. 10B shows the influence of added 5%
noise for f=500 MHz. Low saturation values exhibit greater
sensitivity to the introduced noise than high saturation
values.
The high sensitivity at low saturation values is expected for the
high frequency approximation (Eq. 11). While the absorption
coefficient for an isobestic wavelength does not change with
saturation, lower saturation values yield lower values of the
absorption coefficient for a contrabestic oxyhemoglobin wavelength;
this yields lower values of .theta..sup..lambda.2 -.theta..sub.0 in
the denominator of Eq. 11. Thus, the .mu..sub.a ratio, at the two
wavelengths, is more sensitive to noise at low saturation
values.
FIGS. 11A and 11B show simulation results obtained using the low
frequency approximation (2.pi.f<<.mu..sub.a c) for modulation
frequencies f=10, 50 and 200 MHz, assuming
.theta..sub.0.sup..lambda.1 =.theta..sub.0.sup..lambda.2
=.theta..sub.0, and .mu..sub.a
c.apprxeq.2.multidot.10.sup.9.multidot..theta..sub.0. As shown in
FIG. 11A, the low frequency approximation introduces lower error
for the "intermediate" frequency of 200 MHz than the high frequency
approximation shown in FIG. 10A. However, the low frequency
approximation is much more sensitive to noise as shown in FIG. 11B.
The relatively high sensitivity is again expected because the ratio
of the absorption coefficients at the two wavelengths is obtained
from the square the phase shift ratio, i.e.,
.mu..sub.a.sup..lambda.2 /.mu..sub.a.sup..lambda.1
=(.theta..sup..lambda.1 /.theta..sup..lambda.2).sup.2.
Thus, when using the high and low frequency approximation, the
calculated data may need to be corrected. The correction can be
made by using look-up tables or other methods, such as dual
frequency phase modulation measurement (Eq. 14) or phase modulation
measurements with dual source-detector separation, to obtain more
accurate information about the background phase shift.
FIG. 12 shows simulation results for the oxygen saturation obtained
using Eq. 4 to calculate the ratio of absorption coefficients at
the two wavelengths. This simulation assumed a correct value of the
effective scattering coefficient (.mu..sub.s '=7 cm.sup.-1) and
varied the "selected" tissue saturation (and thus the tissue
absorption). For each "selected" saturation, the simulation
calculated the absorption coefficient solving Eq. 4, while
numerically varying .mu..sub.s ' from 3 cm.sup.-1 to 13 cm.sup.-1
using the Newton-Raphson method. For each .mu..sub.s ', the error
in the calculated saturation Y was calculated by subtracting the
"selected" saturation from the "back-calculated" saturation. As
shown in FIG. 12, for example, for a error of 3 cm.sup.-1 in
.mu..sub.s, the mean error in Y is about 2.5%, while the standard
deviation does not exceed 1.59%. Thus, by employing Eq. 4, the
phase modulation system can use an approximate value of the
effective scattering coefficient to measure the oxygen saturation.
The oxygen saturation is quite insensitive to the selection of the
effective scattering coefficient as the introduced error is reduced
by taking the ratio of the absorption coefficients.
The phase modulation system is calibrated initially and may be
recalibrated after several measurements to obtain a correct phase
reading and an average drift. Another type of a phase modulation
system is PMD-3000 (available from NIM Incorporated, Philadelphia,
Pa.), which is also described in U.S. Pat. No. 5,122,974. This
phase modulation system uses two laser diodes at 754 nm and 780 nm,
each having an average signal power 5 mW. The two wavelengths are
time shared using a mechanical shutter before the light is
introduced in the tissue and then detected by a Hamamatsu R928 PMT
detector. The system uses two frequencies of 200.000 MHz and
200.025 MHz, and the detected signal is demodulated by heterodyning
the second dynode of the PMT detector. The detected amplitude is
used in a feed-back loop as an automatic gain control.
The phase detector of the system provides a voltage output that is
converted then to the phase as specified by the manufacturer. There
are several techniques to determine the voltage-to-phase conversion
curve, which ideally should be linear and the precision should be
better that 0.1.degree.. The conversion curve can be verified by
changing the pathlength of the electrical or optical signal by
changing the physical length of an electrical line. Here, one has
to watch for a line mismatch that can potentially create
measurement problems. Alternatively, the conversion curve can be
verified by changing the source detector separation on an optical
bench and measuring the corresponding voltage difference at the
output of the phase detector. One has to prevent the phase
amplitude cross-talk and operate the system at a proper
signal-to-noise level.
Alternatively, one can simulate a real experiment by using a tank
containing an Intralipid.TM. solution of known absorption and
scattering properties. (See Sevick et al., Analytical Biochemistry
Vol. 195, p. 341.) The source-detector geometry resembles the
actual tissue measurement geometry. The measured absorption
coefficient can thus be compared to the known absorption
coefficient. The voltage-to-phase curve is calibrated by taking
multiple points at different blood concentrations.
The phase modulation system also has a reference phase
(.theta..sub.instr) that of course affects .theta..sub.0. The
instrumental reference phase can be determined empirically or can
be measured by butt-coupling the source and detector fibers. In
this arrangement, the detected optical signal should be attenuated
with a neutral density or NTR filter so the detector works in the
same signal power range as for the in vivo tissue measurements.
The instrumental reference phase can also be measured using a dual
channel phase modulation system that provides both a phase output
and an amplitude output. In this measurement, the above model
should have similar scattering and no absorption, or known
scattering and absorbing properties. The dual channel phase
modulation system can resolve both .mu..sub.s ' and .mu..sub.a,
which in turn are used to calculate the instrumental reference
phase. Furthermore, the instrumental reference phase can also be
determined by measuring the phase shift at different
source-detector separations.
The phase modulation system can use the amplitude in a feedback
arrangement to control the laser intensity. (This type of feedback
is similar to the automatic gain control (AGC) technique described
above.) The intensity is adjusted in discrete steps so that no
change in the laser intensity occurs during the measurement. This
feedback system can measure tissue at a wide range of
source-detector separations or background absorptions; there is no
need to select an optical attenuator or adjust the gain (high
voltage) of the detector. Furthermore, the detector can be operated
in the optimum high voltage for all measurements.
In an experimental study, six newborn piglets, age one to five
days, were used (average weight--2.0 kg). After anesthesia and
surgery, they were randomized either to preexisting mixed acidosis
with a pH less than 7.00 and a PCO.sub.2 larger than 8.0 kPa, or a
normal pH and pCO.sub.2. The acidosis was induced by infusing
lactic acid in a vein, and CO.sub.2 was added to the inspired air.
Once the piglets were stabilized, the fraction of oxygen in the
inspired air (the FiO.sub.2) was reduced from 21% to 6% for 30-40
minutes and then the piglets were resuscitated. Mean arterial blood
pressure was kept above 40 mmHg at all times using an intravenous
adrenaline infusion.
A PMD-3000 system was used to perform the phase modulation
measurements. Part of the scull skin was removed and the optical
probes were fixed directly to the scull. Typical separations used
were 1.7-2 cm. FIGS. 13A and 13B depict the filtered raw data and
saturation calculation from a typical measurement. The filtering
was done digitally by applying a median filter (kernel size 5)
twice followed by a smoothing filter (kernel size 11). The
saturation was calculated by numerically solving Eq. 4 for the two
wavelengths in order to compute the .mu..sub.a ratio as discussed
above. The .mu..sub.s ' value for the pigs was selected to be 12
cm.sup.-1.
During the experimental study, the venous and arterial blood was
sampled regularly and blood saturation was immediately calculated.
Cerebro-venous saturation values were obtained through an
indwelling superior sagittal sinus line and arterial values from a
catheter in the femoral artery. The influence of the arterial blood
sampling can been seen on FIG. 13B, where the observable sampling
points have been marked with arrows, and the local variations are
due to the local blood volume changes. The characteristic values of
hemoglobin saturation for venous (Hbv) and arterial (Hba) blood are
given in FIG. 13B as individual points.
The calculated saturation is somewhat higher than what was expected
for the 6% FiO.sub.2 interval and lower for the 21% interval. This
discrepancy can be correlated by measuring or compensating for
water absorption, geometry and scull influence. Furthermore, the
extinction coefficients were linearly interpolated for the used
wavelengths from charts, and there are random errors introduced in
the measurement or derivation of the .theta..sub.instr.sup.754 and
.theta..sub.instr.sup.780 which may lead to systematic errors in
the calculation.
Additional embodiments are within the following claims:
* * * * *